Working electrode of a continuous biological sensor

ABSTRACT

A working electrode for a subcutaneous sensor for use with a continuous biological monitor for a patient is disclosed. The working electrode includes a conductive substrate and a carbon-enzyme layer on the conductive substrate. The carbon-enzyme layer includes a polyurethane or silicone crosslinked with an acrylic polyol, and an enzyme fully entrapped by the polyurethane or silicone crosslinked with the acrylic polyol. The enzyme is selected according to a biological function to be monitored. The carbon-enzyme layer also includes a carbon material. The carbon-enzyme layer is electrically conductive and facilitates a generation of either peroxide or electrons within the carbon-enzyme layer responsive to reacting the enzyme with a target biologic from blood of the patient.

RELATED APPLICATIONS

This application is a divisional of U.S. Non-Provisional patentapplication Ser. No. 16/375,884, filed Apr. 5, 2019, and entitled“Enhanced Carbon-Enzyme Membrane for a Working Electrode of a ContinuousBiological Sensor”; which claims priority to (1) U.S. ProvisionalApplication No. 62/653,821, filed Apr. 6, 2018, and entitled “ContinuousGlucose Monitoring Device”; (2) U.S. Provisional Application No.62/796,832, filed Jan. 25, 2019, and entitled “Carbon Working Electrodefor a Continuous Biological Sensor”; and (3) U.S. ProvisionalApplication No. 62/796,842, filed Jan. 25, 2019, and entitled “EnhancedMembrane Layers for the Working Electrode of a Continuous BiologicalSensor”; each of which is incorporated herein by reference as if setforth in their entirety.

BACKGROUND

Monitoring of glucose levels is critical for diabetes patients.Continuous glucose monitoring (CGM) sensors are a type of device inwhich glucose is measured from fluid sampled in an area just under theskin multiple times a day. CGM devices typically involve a small housingin which the electronics are located and which is adhered to thepatient's skin to be worn for a period of time. A small needle withinthe device delivers the subcutaneous sensor which is oftenelectrochemical.

Glucose readings taken by the sensor can be tracked and analyzed by amonitoring device, such as by scanning the sensor with a customizedreceiver or by transmitting signals to a smartphone or other device thathas an associated software application. Software features that have beenincluded in CGM systems include viewing glucose levels over time,indicating glucose trends, and alerting the patient of high and lowglucose levels.

Medical patients often have diseases or conditions that require themeasurement and reporting of biological conditions. For example, if apatient has diabetes, it is important that the patient have an accurateunderstanding of the level of glucose in their system. Traditionally,diabetes patients have monitored their glucose levels by sticking theirfinger with a small lance, allowing a drop of blood to form, and thendipping a test strip into the blood. The test strip is positioned in ahandheld monitor that performs an analysis on the blood and visuallyreports the measured glucose level to the patient. Based upon thisreported level, the patient makes critical health decisions on what foodto consume, or how much insulin to inject. Although it would beadvantageous for the patient to check glucose levels many timesthroughout the day, many patients fail to adequately monitor theirglucose levels due to the pain and inconvenience. As a result, thepatient may eat improperly or inject either too much or too littleinsulin. Either way, the patient has a reduced quality of life andincreased risk of doing permanent damage to their health and body.Diabetes is a devastating disease that if not properly controlled canlead to terrible physiological conditions such as kidney failure, skinulcers, or bleeding in the eyes and eventually blindness, pain and oftenthe amputation of limbs.

Complicating a patient's glucose monitoring, it is known that bloodglucose levels can significantly raise or lower quickly, due to severalknown and unknown causes. Accordingly, a single glucose measurementprovides only a brief snapshot of the instantaneous glucose level in apatient's body. Such a single measurement provides little informationabout how the patient's use of glucose is changing over time, or how thepatient reacts to specific dosages of insulin. Accordingly, even apatient that is adhering to a strict schedule of finger pricking andstrip testing, the patient will likely be making incorrect decisions asto diet, exercise, and insulin injection. Of course, this is exacerbatedby a patient that is less consistent on their strip testing. To give thepatient a more complete understanding of their diabetic condition and toget a better therapeutic result, some diabetic patients are now usingcontinuous glucose monitoring.

The CGM sensor is typically temporarily adhered to the patient's skinwith an adhesive pad, and the CGM sensor couples to a small housing inwhich electronics are located. The CGM sensor typically has a disposableapplicator device that uses a small introducer needle to deliver the CGMsensor subcutaneously for the patient. Once the CGM sensor is in place,the applicator is discarded, and the electronics housing is attached tothe sensor. Although the electronics housing is reusable and may be usedfor extended periods, the CGM sensor and applicator need to be replacedoften, usually every few days.

It will be understood that, depending upon the patient's specificmedical needs, that continuous glucose monitoring may be performed atdifferent intervals. For example, some continuous glucose monitors maybe set to take multiple readings per minute, whereas in other cases thecontinuous glucose monitor can be set to take readings every hour or so.It will be understood that a continuous glucose monitor may sense andreport glucose readings at different intervals, and the reading rate maychange depending on past measurements, time of day, or other criteria.

Electrochemical glucose sensors operate by using electrodes whichtypically detect an amperometric signal caused by oxidation of enzymesduring conversion of glucose to gluconolactone. The amperometric signalcan then be correlated to a glucose concentration. Two-electrode (alsoreferred to as two-pole) designs use a working electrode and a referenceelectrode, where the reference electrode provides a reference againstwhich the working electrode is biased. The reference electrodesessentially complete the electron flow in the electrochemical circuit.Three-electrode (or three-pole) designs have a working electrode, areference electrode and a counter electrode. The counter electrodereplenishes ionic loss at the reference electrode and is part of anionic circuit.

Unfortunately, the current cost of using a continuous glucose monitor isprohibitive for many patients that could benefit greatly from its use.As described generally above, a continuous glucose monitor has two maincomponents. First, a housing for the electronics, processor, memory,wireless communication, and power. The housing is typically reusable,and reusable over extended periods of time, such as months. This housingthen connects or communicates to a disposable CGM sensor that is adheredto the patient's body, which uses an introducer needle to subcutaneouslyinsert the sensor into the patient. This sensor must be replaced,sometimes as often as every three days, and likely at least once everyother week. Thus, the cost to purchase new disposable sensors representsa significant financial burden to patients and insurance companies.Because of this, a substantial number of patients that could benefitfrom continuous glucose monitoring are not able to use such systems andare forced to rely on the less reliable and painful finger stickmonitoring.

SUMMARY

In some embodiments, a continuous glucose monitoring sensor includes aworking electrode, a reference electrode and a counter electrode. Theworking electrode has a first wire with a first flat surface and anelectrochemical element on the first flat surface. The referenceelectrode has a second wire with a second flat surface, and the counterelectrode has a third wire with a third flat surface. The first wire,the second wire and the third wire serve as sensor wires for the workingelectrode, the reference electrode and the counter electrode. The secondflat surface and the third flat surface face toward each other.

In another embodiment, a novel working electrode is disclosed for use ina continuous biological sensor. The working electrode uses a plasticsubstrate that is coated with a specially formulated carbon containingcompound. This carbon-containing compound is an aqueous dispersion of acarbon material in an elastomeric material. The carbon-compound isapplied to the plastic substrate, and then further membranes andcoatings are applied to form the working electrode. The workingelectrode may then be associated with one or more reference electrodesor counter electrodes to form the biological sensor.

In one example, the plastic substrate can be polyethylene,polypropylene, polystyrene, polyvinyl chloride, or polylactic acid, andmay be formed into an elongated wire. The carbon material may be, forexample, graphene, diamagnetic graphite, pyrolytic graphite, pyrolyticcarbon, carbon black, carbon paste, or carbon ink, which is aqueouslydispersed in an elastomeric material such as polyurethane, silicone,acrylates or acrylics. Optionally, selected additives may be added tothe carbon compound prior to it being layered onto the plastic wire.These additives may, for example, improve electrical conductivity orsensitivity, or act as a catalyst for target analyte molecules.

In one particular application, the plastic substrate is formed into anelongated wire, and is then coated with a carbon compound that has acarbon material aqueously dispersed in an elastomeric material. Anadditive may be added to the carbon compound that acts as a hydrogenperoxide catalyst, such as Phthalocyanine or Prussian blue. Also, theadditive may be in the form of a metal oxide to enhance electricalcharacteristics, with the preferred metal oxides formed with Copper,Nickel, Rh, or Ir.

Advantageously, a working electrode may be constructed that is durable,strong, flexible and has exceptional electrical and sensitivitycharacteristics. Further, as the working electrode may be constructedwithout the use of expensive and rare platinum, a much morecost-effective working electrode can be provided. Such a platinum-freeelectrode will enable less expensive sensors to be provided to patients,thereby allowing more patients to obtain the substantial benefits ofcontinuous monitoring, and in particular, continuous glucose monitoring.This also allows more flexibility in the mechanical design andconstruction of sensors. Additionally, this design allows for otheranalytes/enzymes beyond glucose where many enzymes systems requirecarbon based electrodes for best performance.

In yet another embodiment, a sensor for a continuous biological monitoris disclosed that has a working electrode with (1) a new interferencelayer for enhancing and stabilizing the interaction of hydrogen peroxidewith a conductor layer and (2) an enhanced glucose limiting layer thatis formed of physical hydrogen bonds. Although these inventive aspectsmay be used independently, they combined to form a highly desirable newworking electrode and sensor. The new sensor is easier and lessexpensive to manufacture than prior devices, and provides improvedsensitivity, better linearity and enhanced accuracy. As compared toprior working sensors, the new interference layer more preciselyregulates the flow of hydrogen peroxide from an enzyme membrane to itselectrical conductor, and enables greater interaction between hydrogenperoxide and the surface of the electrical conductor. The new sensoralso has an outer protective glucose limiting layer that is formed usingphysical hydrogen bonds instead of providing for chemical cross-linking.

In one example of the interference layer, an interference compound iselectrodeposited onto a conductive substrate, and the enzyme layer isapplied over the interference compound. The interference compound is 1)nonconducting, 2) ion passing, and 3) permselective according tomolecular weight. Further, it is electrodeposited in a thin andconformal way, enabling more precise control over the flow of hydrogenperoxide from the enzyme layer to the conductive substrate. In oneparticular example, the interference material is made by mixing amonomer with a mildly basic buffer, and then electropolymerizing themixture into a polymer. For example, the monomer may be 2-Aminophenol,3-Aminophenol, 4-Aminophenol, Aniline, Naphthol, Phenylenediamine, orblends thereof which are mixed with a buffer and electropolymerized intoa polymer. It will be appreciated that other monomers may be used. In amore specific example, the monomer is 2-Aminophenol and the buffer isPhosphate Buffered Saline (PBS) at about 8 pH. The monomer and thebuffer are mixed and electropolymerized into the polymerPoly-Ortho-Aminophenol (PoAP). The PoAp is then electrodeposited ontothe conductive substrate. The permselectivity of the PoAP may beadjusted by the pH of the buffer, for example by adding sodium hydroxide(NaOH).

In one example of the glucose limiting layer, 1) a hydrophilic bondingmaterial, 2) a hydrophobic bonding material, and 3) a solvent are mixedtogether to form a bonding gel. The bonding gel is then applied over theenzyme membrane layer, and the gel is cured. The hydrophilic material istypically selected to be a high molecular weight, readily dispensable,and provide for strong hydrogen bonding. In one particular example, thehydrophilic bond material is Polyvinylpyrrolidone (PVP). The hydrophobicmaterial is selected to be biocompatible, and to have sufficienthardness and still provide for appropriate interaction with thehydrophilic material and the solvent. It is been found that polyurethaneand silicone are desirable hydrophobic materials. Finally, the solventis selected to be polar, binary and volatile enough to support thecuring requirements.

Advantageously, the novel interference layer and the novel glucoselimiting layer are both economically manufacturable to provide morecost-effective working electrodes. Further, both new membranes providefor enhanced linearity and overall detection characteristics for theworking electrode. In one example, the interference layer isnon-electron conducting, ion passing, and is permselective for molecularweight and the glucose limiting layer is a self cross linkingformulation of acrylic poly and polyurethane.

In yet another embodiment of the invention, the working wire has anenzyme layer comprising an aqueous emulsion of a polyurethane and GOxblend, which is applied to the working wire and cured. The new enzymelayer has better stability and full entrapment of GOx, a more evendispersion, and enables higher loading of GOx and better overall sensorsensitivity. It will also be understood that for measuring othermetabolic functions other enzymes could be substituted for GOx.

In another embodiment of the invention, the working wire has acarbon-enzyme layer comprising an aqueous emulsion of a polyurethane,carbon and GOx blend, which is applied and cured to a plastic substratefor a working wire. The new carbon-enzyme layer has better stability andfull entrapment of GOx, a more even dispersion, and enables higherloading of GOx and better overall sensor sensitivity. Further, thecarbon-enzyme layer is able to directly generate free electrons inproportion to the amount of glucose reacted, thereby by eliminating anyneed for expensive platinum. It will also be understood that formeasuring other metabolic functions other enzymes could be substitutedfor GOx.

BRIEF DESCRIPTION OF THE DRAWINGS

These and other objects and advantages of the present disclosure willbecome apparent upon reading the following detailed description and uponreferring to the drawings and claims.

FIG. 1 shows various views of flat-surface electrodes, in accordancewith some embodiments.

FIGS. 2A-2D show various views of flat-surface electrodes with atriangular support core wire, in accordance with some embodiments.

FIG. 3 shows an electrochemical element being mounted to a flat-surfaceelectrode, in accordance with some embodiments.

FIG. 4A is a not-to-scale illustration of a carbon coated wire for aworking electrode in accordance with some embodiments.

FIG. 4B is a not-to-scale illustration of a carbon coated wire for aworking electrode in accordance with some embodiments.

FIG. 4C is a not-to-scale illustration of a carbon coated wire for aworking electrode in accordance with some embodiments

FIG. 5 is a flowchart of general manufacturing steps for making aworking electrode in accordance with some embodiments.

FIG. 6 is a not-to-scale cross-sectional block diagram of a prior artsingle wire sensor.

FIG. 7A is a not-to-scale cross-sectional block diagram of a 2-wiresensor having an interference membrane layer in accordance with someembodiments.

FIG. 7B is a not-to-scale cross-sectional block diagram of a 2-wiresensor having an interference membrane layer and a coated referenceelectrode in accordance with some embodiments.

FIG. 8 is a flowchart of general manufacturing steps for making a 2-wiresensor having an interference membrane layer in accordance with someembodiments.

FIG. 9A is a not-to-scale cross-sectional block diagram of a 2-wiresensor having an interference membrane layer and a glucose limitinglayer in accordance with some embodiments.

FIG. 9B is a not-to-scale cross-sectional block diagram of a 2-wiresensor having an interference membrane layer, a glucose limiting layer,and a coated reference electrode in accordance with some embodiments.

FIG. 10 is a flowchart of general manufacturing steps for making andapplying a glucose limiting layer in accordance with some embodiments.

FIG. 11 is a flowchart of general manufacturing steps for making andapplying an enzyme layer in accordance with some embodiments.

FIG. 12 is a not-to-scale cross-sectional block diagram of a 2-wiresensor having an enzyme layer, an interference layer, a glucose limitinglayer, and a non-platinum substrate in accordance with some embodiments.

FIG. 13A is a not-to-scale cross-sectional block diagram of a 2-wiresensor having a carbon/GOx membrane for direct generation of electronsor peroxide and an interference layer in accordance with someembodiments.

FIG. 13B is a not-to-scale cross-sectional block diagram of a 1-wiresensor having a carbon/GOx membrane for direct generation of electronsor peroxide and an interference layer in accordance with someembodiments.

FIG. 14A is a not-to-scale cross-sectional block diagram of a 2-wiresensor having a carbon/GOx membrane for direct generation of electronsor peroxide in accordance with some embodiments.

FIG. 14B is a not-to-scale cross-sectional block diagram of a 1-wiresensor having a carbon/GOx membrane for direct generation of electronsor peroxide and in accordance with some embodiments.

FIG. 15 is a not-to-scale cross-sectional block diagram of 1-wiresensors having a working wire with a carbon/GOx membrane for directgeneration of electrons or peroxide and an attached reference wire inaccordance with some embodiments.

FIG. 16 is a flowchart of general manufacturing steps for making andapplying a carbon/GOx membrane for direct generation of electrons orperoxide in accordance with some embodiments.

DETAILED DESCRIPTION

The present disclosure relates to structures and processes for sensorsused in a continuous metabolic monitor, such as a continuous glucosemonitor. In particular, the present devices and methods describe novelmembranes and substrates for use with a working electrode in acontinuous metabolic sensor. Cost can be a prohibiting factor forpatients who could benefit from the use of CGMs. Accordingly, there is asignificant need in the market for a lower-cost sensor for continuousbiological monitors. It will be understood that cost reduction may beobtained by reducing the manufacturing cost of the sensor itself, byincreasing the length of time between sensor replacements, or by acombination of both reducing cost and increasing the useful life. Bydecreasing the cost of sensors for continuous monitoring, more patientscould benefit from the increased quality of life and enhancedtherapeutic effect of continuous monitoring

Most CGM sensor designs are either planar (flat substrate) orwire-based. Planar types are more amenable to use with 3-poleelectrochemical designs since simple wire traces and small electrodescan be easily constructed. However, planar types have deficienciesregarding physiology since a planar substrate has some directionalityand also has sharp edges due to its geometry, which leads to a moreaggressive biologic response to the device. Wire-based systems result inbetter physiological responses from the patient than planar systems dueto the smooth nature of their geometry but have been mostly confined toa single wire for ease of insertion through a needle. This single wireconstraint due to the space limitations of needle-based sensor deliverytypically limits the designs to 2-pole electrochemical designs. The2-pole design has an added drawback of making the reference electrodenon-renewable and thus the electrode material is consumed to completethe electrochemical circuit, which limits the working life of thesystem.

A challenge of wire-based sensor designs is making electricalconnections on the distal end. The single wire configuration requiresin-situ fabrication of working membranes and chemistries and thus limitsthe approaches and materials that can be used in such designs. Separatewires for working, reference and counter electrodes would be ideal forease of fabrication; however, this approach is limited by the internaldiameter of the insertion needles.

The present embodiments disclose a wire-based 3-pole electrochemicaldesign that solves deficiencies of the aforementioned designs. Theworking chemistries are made separately from the wires and then bondedto the underlying sensor wires. This allows for lower cost materials andmethods since components of the present CGM devices can be madeindependently from each other. Also, more cost-effective scaledmanufacturing is enabled since manufacturing the wires separately doesnot require 100% sensor quality testing, and quality testing can beperformed on a sheet or lot basis. Some embodiments of the disclosedwire-based systems use carbon-based, such as graphene-based, electrodesmanufactured in large-scale sheets with working chemistries that arethen attached to the working electrode.

Wire-based 3-pole Electrode Design

FIG. 1 illustrates an embodiment of a wire-based 3-pole system 100 of acontinuous glucose monitoring sensor in which a split wire design isused. In this embodiment, a portion of a wire such as a half-wire isprovided for a reference electrode 110 and a counter electrode 120, eachhaving a flat surface across approximately its diameter such that thewires have semi-circular cross-sections. In some embodiments, the flatsurface of the reference electrode 110 and the flat surface of thecounter electrode 120 face toward each other. Each half-wire electrode,such as the reference electrode 110 and the counter electrode 120, mayhave a partial surface area such as 82% of the surface area of a fullwire having the same diameter, while still allowing the reference andcounter electrode assembly to fit within a small diameter insertionneedle 102 for insertion under the skin. In other words, the split-wireconfiguration enables the reference electrode 110 and the counterelectrode 120 to provide nearly the same surface area as two full wireelectrodes, but only occupy the space of one wire within the insertionneedle 102 instead of two full wires. Although half-wires are depictedfor the reference electrode 110 and counter electrode 120—where eachwire has been split along its diameter along a length of the wire—otherpartial fractions of the wire may be utilized to form the flat surfaceelectrodes such as, for example, 30% to 70%, or 40% to 60%, typicallydefined by a chord drawn across the circular cross-section of the wireto get larger percentages of the surface area of full wire.

A working electrode is fabricated by also creating a flat portion on awire. FIG. 1 shows two embodiments—a 1-sided working electrode 130 and a2-sided working electrode 135—either of which may be used. The 1-sidedworking electrode 130 has a semicircular cross-section where half of thewire's cross-sectional area has been removed, while the 2-sided workingelectrode 135 has a rectangular cross-section where portions of the wireabove and below the flat portion have been removed. The portions removedmay be equal or one portion—either the top portion or the bottomportion—may be larger than the other. The flat portion(s) of workingelectrode 130 or 135 is used to support an electrochemical element whichis the reactive component that senses glucose in the patient'sinterstitial fluid.

FIG. 1 also shows insertion of the electrodes into insertion needle 102,where it can be seen that this 3-pole design of the sensor occupies aspace within the needle lumen equivalent to only two wires instead ofthree wires. The working electrode (where 2-sided working electrode 135is shown in this illustration) utilizes the space of one wire, and thereference electrode 110 and counter electrode 120 together occupy thespace of another wire. Diameters of the wires used for the referenceelectrode 110, counter electrode 120 or working electrode 135 may be,for example, from 0.002 inches to 0.007 inches. The length or surfacearea of the electrode portions themselves can be tailored according tothe desired sensor sensitivity and required design specifications.

Other embodiments of systems in which flat surface electrodes are usedin a continuous glucose monitoring sensor are shown in FIGS. 2A-2D. Inthe radial cross-sectional views of designs 200 (FIG. 2A) and 210 (FIG.2B), compact systems are assembled from a support core wire 140 having atriangular cross-section surrounded by working electrode 135, referenceelectrode 110 and counter electrode 120 facing the flat surfaces of thetriangular core wire 140. Each of the wires for the working electrode135, reference electrode 110 and counter electrode 120 have flatsurfaces that are positioned to be facing a surface of thetriangular-shaped core wire 140. As shown in FIGS. 2A-2B, the referenceelectrode 110 and counter electrode 120 are approximately semi-circularin cross-section, while the working electrode 135 can be semicircular(design 200 of FIG. 2A) or rectangular (design 210 of FIG. 2B) incross-section. Schematics 220 (FIG. 2C) and 230 (FIG. 2D) provide alongitudinal cross-sectional view and a perspective view, respectively,of the end of the triangular support wire, showing that this triangulardesign can be a completely self-inserting sensor. That is, a tip 142 ofthe triangular core wire 140 may be sharpened to a point, or pointed,such that the sensor can be inserted directly, without requiring the useof a needle to place the sensor within the subcutaneous tissue.

The flat surfaces of the electrodes in these various embodiments providesupport for fragile electrochemistry materials, such as a carbon-basedsheet which is typically brittle. In one example, a support sheet (e.g.,made of pyrrole or polyaniline) can be created, and then a carbonmaterial is deposited onto the support sheet. The support sheet providesa substrate to which the carbon bonds well, and also should beconductive to electrically couple the electrochemical (e.g.,carbon/pyrrole) sheet to the electrode wire. The conductive sheetmaterial can then be impregnated or coated with sensing chemistries viavarious drawn membrane or spin coating techniques.

The electrochemistry material sheet can be made separately from theelectrode wire, and then mounted on the flat surface of the electrode asshown in FIG. 3. In this example of FIG. 3, a wire 310 having insulation320 surrounding a conductive wire core 330 has a portion of its endremoved to form a flat surface 340. A carbon or carbon/graphene/pyrrolesheet 350 is cut to size and placed on the flat surface 340 of the flatelectrode wire 310. For example, once the flat sheets are fabricatedwith sensing chemistries, these sheets 350 can be laser cut into smallportions and then assembled onto the flat surface 340 of the wire 310.

The support sheet can be made by, for example, depositing a pyrrolelayer to make electrical contact to the flat surface of the electrode.In other embodiments, an electro-polymerization of additional pyrrolecan be used to connect the electrode metal to the sheet, or conductiveadhesives or other electrical contact bonding methods can be used tomake electrical contact as well.

In other embodiments, the electrochemistry components can be formedin-situ on the electrode instead of forming a sheet separately from theelectrode. For example, an alternative fabrication method for in-situcreation of sensing chemistry and membrane may include pad or screenprinting, painting, or 3D-printing directly onto the flat plane(s) ofthe wire.

The carbon material can be in the form of, for example, an ink or apaste, and the carbon can include various allotropes such as but notlimited to graphite, graphene, fullerenes, and/or nanotubes. Materialsother than pure carbon can be used, including platinum black, carbonplatinum pastes, carbon gold pastes or other known working electrodesurface materials, alone or in combination (e.g., carbon, platinum,gold, palladium, rhodium, iridium). In some embodiments, high surfacearea nano-porous materials of graphene and/or other nanomaterials can beused, to increase the number of active chemical sites available forreactions.

Carbons are lower cost than the metals that are typically used forbiocompatible applications (e.g., gold and platinum). However, due tothe inherent brittle nature of carbon materials, carbon-based electrodeshave been conventionally used in planar style electrodes (such as fingersticks) where the carbon can be supported by the planar substratewithout applying undue mechanical loads on the electrode. The presentembodiments overcome the difficulties of using carbon-based materials ona wire electrode by providing the mechanical support required for thecarbon material and by eliminating the typical need for in-situfabrication of the working chemistries on the wire (although in-situfabrication may be used).

After the sensing chemistry has been created on the electrode, whetherseparately or in-situ, a final dip coating may be used to seal theentire system using hoop strength created by polymer shrinkage upondrying. This final polymer layer also serves as a biocompatible andglucose limiting membrane required for creating a linear glucoseresponse, and provides the biosafety required for an implanted sensor.

The present flat-wire embodiments may also be used to optimize theelectrochemical substrate so that it can be tuned for direct electrontransfer chemistries by keeping the redox center close to the porouscarbon surface or within encapsulating polymers. One such embodimentuses an aminophenol covalently bonded to the carbon electrode byelectrografting and is subsequently linked by diazonium chemistry toglucose oxidase (GOx) to provide direct electron transfer. Embodimentscan be directly used with conductive polymers (e.g., PEDOT-PSS,poly-pyrroles, polyanilines, naphthol, phenylenediamine, etc.) formedin-situ on a porous carbon sheet that would work with normal enzymes(either glucose oxidase (GOx) or glucose dehydrogenase (GDH)) and/orenzymes with a mediator to create hybrid enzyme systems that alter theneed for high bias voltages and thus reduce interferences from allsources.

In some embodiments, a redox enzyme can be immobilized on the electrodesurface in a new manner such that direct electron transfer between theactive side of the enzyme and the transducer is possible. The majorunique character of such embodiments of an amperometric glucose sensoris that its biased potential is in the range of 0 to −0.5V, ideally tobe around −0.1V. In comparison, a conventional CGM sensor has a biasedpotential of typically +0.55V. There are two major methods to achievethe lower biased potential of the present designs. A first method is anin-situ electro-polymerization of a conductive polymer with a redoxenzyme. The sensing layer is formed by applying potential cycles orsequences of suitable potential pulses with the enzyme andmonomer/comonomers solution. An advantage of this approach is that thefilms are formed exclusively on the electrode surfaces due to theelectrochemical initiation of the deposition process. A second method isthe incorporation of a redox mediator into the polymers or theprepolymer. The polymer that contains the redox mediator can bephysically mixed with the enzyme, then be deposited onto the electrodesthrough dip coating, spin coating or other coating methods. This canalso be achieved through the in-situ polymerization of theredox-mediator-containing prepolymer with other active prepolymers inthe presence of the enzyme solution and the electrodes. The resultingsensing layer on the electrode contains the matrixed enzyme inside thepolymer network with the covalently linked redox mediator.

Carbon Substrate

In some embodiments, a cost-effective platinum-free sensor is used in acontinuous biological monitoring system. Embodiments provide for asubstantial reduction in cost for the manufacture of the workingelectrode for such a biological sensor. Although the embodiments arediscussed primarily for the use in continuous glucose monitoring, itwill be understood that many other uses for biological sensing existthat would benefit from a reduced cost sensor and working electrode.

Typically, a sensor for a continuous biological monitoring system has aworking electrode and a reference electrode. The working electrode andreference electrode are constructed and arranged such that they cansense the concentration of an analyte in the patient, oftentimes bymeasuring a concentration or ion flow within the blood or other bodyfluids, such as interstitial fluid (ISF). It will be understood that asensor may include multiple working wires, multiple referenceelectrodes, and counter electrodes.

Generally a working electrode needs to be constructed to meet threebasic requirements. First, it must be strong enough to withstandinsertion under the patient's skin and to withstand the vibrations,shocks, and motions during use. Second, it needs to be flexible enoughto follow a curved path into the skin, and to allow for some movementafter insertion for patient comfort. And third, it needs to provide theelectrical characteristics to support consistent and accurate sensing.Accordingly, known working electrodes typically use some form of aplatinum wire, either a solid platinum wire, or a less expensive metalmaterial (such as tantalum) coated with platinum. It is this relianceand use of platinum that drives some of the high cost of currentbiological sensors.

Advantageously, embodiments of the present disclosure eliminate the needfor expensive and rare platinum to make a working electrode that notonly has sufficient mechanical strength and flexibility but has superiorelectrical and sensing characteristics. Further, embodiments of thepresent working electrodes are constructed of materials known to be safein a human body. This also allows for the use of alternative geometriesof sensors and different styles of sensor manufacturing.

In one particularly cost-effective embodiment, the working electrodeuses a plastic material as a substrate. The plastic material issufficiently strong to support insertion into human body, while havingthe needed flexibility for insertion and patient comfort. This plasticsubstrate can be formed into an elongated wire in many shapes to supportconstruction of different types of sensors. The plastic wire may then becoated with a specially formed carbon compound. Plastic wire has theadded advantage of improved fatigue performance in comparison to ametallic wire of the same dimension. Traditionally, elemental carbonpaste electrodes could not be considered for use on a flexible workingelectrode, as carbon is highly brittle and needed to be rigidlysupported. Also, carbon paste electrodes are usually water-soluble, andtherefore dissolve and degrade when inserted into a wet environment. Andfinally, carbon has a high electrical resistance compared to platinummetal, and therefore is not practically usable as a conductor in abiological sensor. However, the new forms of carbons in a carriercompound used over the plastic wire as disclosed herein overcome theseveral disadvantages of the elemental carbon.

In some embodiments, a carbon compound is prepared as a coating that isan aqueous dispersion of a carbon material with an elastomeric material.For example, the carbon material may be in the form of graphene,diamagnetic graphite, pyrolytic graphite, pyrolytic carbon, carbonblack, carbon paste, or carbon ink. In some cases, to support particularapplications, other additives may be added to the carbon compound forenhanced electrical and response characteristics. For example, ahydrogen peroxide catalyst could be added to the carbon compound tosupport enhanced glucose level sensitivity. It will be understood thatother sensing molecules may be used for other sensing applications.

The carbon compound, as described above, is then applied to the plasticwire. Most often, this would be through a simple dipping process,although it will be understood that the coating could also be sprayed,extruded, deposited, or even printed onto the plastic wire or direct 3-Dprinted substrates. The coated wire may then be processed into a workingelectrode using known processes by adding membranes, associating it witha reference electrode, and adding protective biological coatings.

Referring now to FIG. 4A, FIG. 4B, and FIG. 4C, carbon coated wires 400are illustrated. The carbon-coated wires 400 include wire 411 in FIG.4A, wire 412 in FIG. 4B and wire 413 in FIG. 4C. These illustrations arenot to scale and are used only for descriptive purposes. The carboncoated wires 400 each have a plastic core 415 which is fully surroundedby a carbon compound 418. It will be understood that the plastic core415 may be formed into many different elongated physical shapes. Forexample, as illustrated in FIG. 4A, the plastic core 415 may have acircular cross-section. As illustrated in FIG. 4B, the plastic core 415may have a rectangular or square cross-section. And as illustrated inFIG. 4C, the plastic core 415 may have a triangular cross-section. Itwill be appreciated that many other cross-section shapes may be used.

The carbon compound 418 is formulated to have superior electricalcharacteristics appropriate mechanical characteristics such as strengthand flexibility, and to be cost-effective. For example, a standardcarbon conductive ink has a resistivity of about 23 Ohm/mm², whilecarbon compound 418 can be formulated to have a much more desirableresistivity such as 1-5 Ohm/mm². In this way, the carbon compound 418has been found to have resistivity that is an order of magnitude lowerthan standard carbon conductive inks, dramatically increasing itsutility and performance as a conductor for working wire. Not only is thecarbon coating 418 far less expensive than platinum, it is also easierand more cost-effective to apply as a coating. For example, the carboncompound 418 may be used with a low-cost dipping, spraying, extrusion,depositing, or printing process. It will be understood that the carboncoated wires 400 will be further processed to add membranes andprotective coatings according to the specific application, and that theywill be associated with one or more reference or counter electrodes. Itwill be understood that the association of the working wire with areference wire may be accomplished in several ways. For example, theworking wire and the reference wire may be placed side-by-side, formedconcentrically, wrapped into a twisted relationship, layered, or formedinto any other known physical relationships for a working wire and itsassociated reference wire.

In one example, the carbon coating may be formulated as follows. It willbe appreciated that many other formulations fall within the teachingsherein.

Formulation (% by weight for a total of 100%)

-   -   an aqueous dispersion, comprising        -   40-60% Polyurethane (i.e. Hauthaway HD4661,        -   40-60% Acrylic Polyol (i.e. Acquathane),        -   0.5-5%% Polyvinylpyrrolidone).    -   0.1-0.5% Carbon Black    -   0.05%-0.5% Graphene    -   0.1-0.5% Pyrolytic Graphite    -   0-10% additional water

Referring now to FIG. 5, a process 500 for making a carbon workingelectrode is illustrated. Process 500 begins with selecting a plasticsubstrate material in step 522. This plastic substrate material isselected to have sufficient strength for being inserted under the skinof a patient, as well as flexibility for patient comfort and ease ofmanufacturing. Further, it will be understood that the plastic substrateshould be biologically safe and generally electrically unreactive. Itwill be understood that a wide range of materials meet the mechanicaland functional requirements for the selected plastic substrate. Forexample, numerous organic polymers and thermoplastics may be used. Forillustrative purposes only, the following specific plastic substratematerials may be used: polyethylene, polypropylene, polystyrene,polyvinyl chloride, and polylactic acid. It will be appreciated that awide variety of materials may be used as the plastic substrate.

The selected plastic substrate material is then formed into an elongatedplastic wire in step 523. It will be understood that the wire may takemany cross-sectional shapes, such as circular, square or triangular.Generally, these wires may be formed using well-known extrusionprocesses. The plastic substrate could also be formed into a ribbonwire, or in some cases manufactured by printing such as 3D printing.

A carbon compound is prepared in step 525 for application to the plasticsubstrate. The carbon compound has a carbon material that is aqueouslydispersed in an elastomeric material. The elastomeric material isselected for its mechanical properties, such as strength andflexibility, while the carbon material is selected for an advantageouselectrical property. There are several acceptable elastomeric materialsthat may provide the desired characteristics, for example: aspolyurethane, silicone, acrylates or acrylics. It will be understoodthat other elastomeric materials may be substituted. Experimentalresults related to the present disclosure have shown that the carboncompound, after the elastomeric material is cured, does not delaminatefrom the plastic wire, in contrast to platinum coated of a Tantalumwire.

The carbon material in the carbon compound is selected for enhancedelectrical characteristics. For example, as briefly discussed above,elemental carbon has much too high an electrical resistance to beeffectively used in a working electrode. However, with the addition ofgraphene, diamagnetic graphite, or pyrolytic carbon, the carbon compoundcan be formulated to have advantageous electrical characteristics.Indeed, the loading of the carbon material with the elastomeric materialcan be adjusted to create a carbon compound with a desired resistance,for example 100 ohms/cm² or less. In this way, a working electrode usingsuch a carbon compound coating can be used such that the sensing systemhas a highly desirable signal-to-noise ratio. Elemental carbon would notbe able to enable such a signal-to-noise ratio due to the highelectrical background from the high resistance of carbon electrodes.

Optionally, an additional catalyst or material may be added to thecarbon compound in step 526 to enhance electrical or sensingcharacteristics. For example, metal oxides may be added to the carboncompound for reducing resistivity thereby enabling a working electrodewith a higher signal-to-noise ratio capability as compared to elementalcarbon. For example, in some embodiments metal oxides of nickel orcopper may be used. In some embodiments, metal oxides of Rh and Ir, whenadded to the carbon compound, can enable the working wire to operatewith a lower bias voltage as compared to a wire formed with platinum. Byoperating at a lower bias voltage, a working wire is enabled to operatewith greater sensitivity, and with a lower power consumption.

In another example of an additive in step 526, a catalyst for hydrogenperoxide may be added to the carbon compound. In one example,phthalocyanine or Prussian blue is added to the carbon compound, therebysubstantially increasing the sensitivity of the working wire to hydrogenperoxide, which is highly advantageous to the overall accuracy andsensitivity for a glucose monitor sensor. It will be understood thatother hydrogen peroxide catalysts may be used. Also, for working wiresintended for biological sensing other than glucose, it will beunderstood that other sensing molecules and molecule catalysts may beused.

The coating can then be applied to the plastic wire substrate in step527. As the carbon compound is inexpensive and easy to work with, theplastic substrate may be dipped into the carbon compound. For otherapplications, the carbon compound may be sprayed onto the plastic wire,may be deposited using well-known deposition processes, co-extrusion, ormay be applied using a printing process, such as pad printing. It couldalso be made to 3-D printed. It will be understood that any appropriateapplication process may be used to coat or deposit the carbon compoundonto the plastic substrate. The carbon compound coating then cures priorto further processing.

Once the carbon coated working wire has cured, it may be processed intoa working electrode in step 528. In this way, membranes and protectivecoatings may be added, and the working wire is associated with one ormore reference or counter electrodes. The processes for addingmembranes, protective coatings, and associating with other electrodes iswell known, so will not be described herein. For example, workingelectrodes and reference electrodes may be set side to side, layered,concentrically formed, or wrapped together. It will also be understoodthat some applications will use multiple working electrodes, multiplereference electrodes, or counter electrodes.

Due to the aqueous nature of the carbon containing compound, enzymes orother sensing molecules and chemistries could be included directly in tothe carbon containing compound, improving efficiency of the electrontransfer and further improving signal to noise ratios by removingadditional layers and diffusional distances. This incorporation ofenzymes and other sensing chemistries into the sensor wire itself wouldalso further simplify manufacture of these sensors.

Referring now to FIG. 6, a prior art single wire sensor 600 for acontinuous biological monitor is illustrated. Those skilled in the artwill recognize that sensor 600 is a high level diagram for instructionalpurposes only, and has left out substantial detail to facilitateimproved understanding. As is understood, such a prior art sensor willintegrate both the function of a working electrode and a function of areference electrode on a single wire. It will be understood that asingle wire electrode may be constructed with multiple working electrodelayers and with multiple reference electrode layers. A single wireelectrode may also use or be supplemented with a counter electrode.Although sensor 600 is illustrated as a wire with concentrically formedlayers, it will be understood that the other physical implementationsmay be used, such as layered, spiraled, flat, and other well-knownphysical relationships.

The prior art sensor 600 has an elongated conducting wire 605, which isoften made of solid platinum or a platinum coating on a less expensivemetal or plastic substrate. It will be appreciated that other types ofconducting wires may be substituted. The conducting wire 605 is wrappedwith an electrically insulating layer 614 a. A band 618 of theinsulating layer 614 a is removed during manufacturing that exposes aportion 617 of the platinum wire, which remains uninsulated. The removalof this band 618 must be done very accurately and precisely, as thisaffects the overall electrical sensitivity of the sensor 600. Forexample, this band 618 may be in the order of 20 μm thick, and must becut to about 40 μm, although other thicknesses and widths may be useddepending upon the overall structure of the sensor 600.

A layer 611 of silver or silver chloride is positioned around theelectrically insulating layer 614, and a second layer of electricallyinsulating material 614 b is disposed around the silver/silver chloridelayer 611. During manufacture, a portion 621 of the silver/silverchloride layer 611 needs to be exposed. Typically, this requires preciseremoval of a small portion of layer 614 b using, for example, a laserablation process. A second removal process may also be used at theconnection end of the sensor 600 to expose a small portion of thesilver/silver chloride layer 611 so that a more convenient electricalconnection may be made. Removal of the insulating layers from thesilver/silver chloride layer 611 is a precision operation, as the layermay be in the order of only 20 μm thick. This expensive removaloperation adds substantial cost and manufacturing risk to making thesingle wire sensor 603.

In operation, the glucose limiting membrane 607 substantially limits theamount of glucose that can reach the enzyme membrane 608. By limitingthe amount of glucose that can reach the enzyme membrane 608, linearityof the overall response is improved. The glucose limiting membrane 607also permits oxygen to travel to the enzyme membrane 608. The keychemical processes for glucose detection occur within the enzymemembrane 608. Typically, the enzyme membrane 608 has one or more glucoseoxidase enzymes (GOx) dispersed within the enzyme membrane 608. When amolecule of glucose and a molecule of oxygen (O₂) are combined in thepresence of the glucose oxidase, a molecule of gluconate and a moleculeof hydrogen peroxide (H₂O₂) are formed. The hydrogen peroxide thengenerally disperses both within the enzyme membrane 608 and into the ionconducting layer 609.

At least some of the hydrogen peroxide travels to the window (band 618)in the electrically insulating layer 614 a, where it comes into contactwith the exposed portion 617 of the platinum wire 605. The platinumsurface facilitates a reaction wherein the hydrogen peroxide reacts toproduce water and hydrogen ions which are released into the ionconducting layer 609, and two electrons are generated. The electrons aredrawn into the platinum wire 605 by a bias voltage placed across theplatinum wire 605 and the silver/silver chloride layer 611. Positiveions from the silver/silver chloride layer 611 are released into the ionconducting layer 609 to complete the electrical circuit. In this way,the magnitude of the electrical current on the platinum wire is intendedto be related to the number of hydrogen peroxide reactions, which isintended to be related to the number of glucose molecules oxidized. Inthis way, a measurement of the electrical current on the platinum wireis intended to be associated with a particular level of glucose in thepatient's blood or ISF.

Unfortunately, since the platinum surface (portion 617) has been exposedduring the manufacturing process, an oxidation layer has formed in thewindow 618. This oxidation layer poisons the electrode and interfereswith the exposed platinum's efficiency in converting the hydrogenperoxide. That is, the actual useful exposed area of the exposed portion617 of the platinum wire is substantially reduced by oxidationcontamination, which also may lead to unpredictable and undesirablesensitivity results. In order to overcome this deficiency, the singlewire sensor 603 must be subjected to sophisticated calibration. Further,the bias voltage between the platinum wire 605 and the silver/silverchloride layer 611 must be set relatively high, for example between0.4-1.0 V. Such a high bias voltage is required to draw the electronsinto the platinum wire, but also acts to attract contaminants from theblood or ISF into the sensor. These contaminants such as acetaminophenand uric acid interfere with the chemical reactions, leading to falseand misleading glucose level readings. The single wire sensor 603 isalso expensive to manufacture, due in part to the precise laser ablationneeded to expose the band 618 in the platinum wire 605, as well asexposing the small 621 portion of the silver/silver chloride layer 611.

Embodiments of the present disclosure are directed to a cost-effectivesensor for use in a continuous biological monitoring system. Embodimentsprovide for a substantial reduction in cost for the manufacture of theworking electrode for such a biological sensor. Although the embodimentsare discussed primarily for the use in continuous glucose monitoring, itwill be understood that many other uses for biological sensors existthat would benefit from a reduced cost sensor and working electrodeswith enhanced functionality.

Generally, a sensor for a continuous biological monitoring system isconstructed as a two-wire continuous biological sensor that has aworking electrode and a reference electrode. The working electrode andreference electrode are constructed and arranged such that they cansense the concentration of an analyte in a patient, oftentimes bymeasuring concentration molecules, such as glucose, within the blood orother body fluid, such as ISF. It will be understood that a sensor mayinclude multiple working wires, multiple reference electrodes, andcounter electrodes.

Interference Layer

Referring now to FIG. 7A, a sensor 700 for a continuous biologicalmonitor is generally illustrated. The sensor 700 has a working electrode703 which cooperates with a reference electrode 705 to provide anelectrochemical reaction that can be used to determine glucose levels ina patient's blood or ISF. Although electrode sensor 700 is illustratedwith one working electrode 703 and one reference electrode 705, it willbe understood that some alternative sensors may use multiple workingelectrodes, multiple reference electrodes, and counter electrodes. Itwill also be understood that sensor 700 may have different physicalrelationships between the working electrode 703 and the referenceelectrode 705. For example, the working electrode 703 and the referenceelectrode 705 may be arranged in layers, spiraled, arrangedconcentrically, or side-by-side. It will be understood that many otherphysical arrangements may be consistent with the disclosures herein.

The working electrode 703 has a conductive portion, which is illustratedfor sensor 700 as conductive wire 710. This conductive wire 710 can befor example, solid platinum, a platinum coating on a less expensivemetal or plastic, or as disclosed above, the conductive wire 710 may bea carbon compound coating on a plastic substrate. It will be understoodthat other electron conductors may be used consistent with thisdisclosure. As with prior art working electrodes, working electrode 703has a glucose limiting layer 707, which may be used to limitcontaminations and the amount of glucose that is received into theenzyme membrane 708.

In operation, the glucose limiting membrane 707 substantially limits theamount of glucose that can reach the enzyme membrane 708, for exampleonly allowing about 1 of 1000 glucose molecules to pass. By strictlylimiting the amount of glucose that can reach the enzyme membrane 708,linearity of the overall response is improved. The glucose limitingmembrane 707 also permits oxygen to travel to the enzyme membrane 708.The key chemical processes for glucose detection occur within the enzymemembrane 708. Typically, the enzyme membrane 708 has one or more glucoseoxidase enzymes (GOx) dispersed within the enzyme membrane 708. When amolecule of glucose and a molecule of oxygen (O₂) are combined in thepresence of the glucose oxidase, a molecule of gluconate and a moleculeof hydrogen peroxide are formed. The hydrogen peroxide then generallydisperses both within the enzyme membrane 708 and into interferencemembrane 709.

The interference membrane 709 is layered between the electricalconducting wire 710 and the enzyme membrane 708 in working electrode703. As will be discussed in more detail below, the interferencemembrane 709 can be uniquely formulated to have a more preciseregulation, compared to conventional insulating layers (e.g. layers 614a/b with Ag/AgCl layer 611 of FIG. 6), of the level of hydrogen peroxidemolecules that are enabled to pass from the enzyme membrane layer 708 toa more expansive surface area of the conductive wire 710. Thisinterference membrane 709 may be electrodeposited onto the electricalconducting wire 710 in a very consistent and conformal way, thusreducing manufacturing costs as well as providing a more controllableand repeatable layer formation. The interference membrane 709 isnonconducting of electrons, but will pass negative ions at a preselectedrate. Further, the interference membrane 709 may be formulated to bepermselective for particular molecules. In one example, the interferencemembrane 709 is formulated and deposited in a way to restrict thepassage of larger molecules, which may act as contaminants to degradethe conducting layer 710, or that may interfere with the electricaldetection and transmission processes.

Advantageously, the interference membrane 709 provides reducedmanufacturing costs as compared to known insulation layers, and isenabled to more precisely regulate the passage of hydrogen peroxidemolecules to a wide surface area of the underlying conductive layer 710.Further, formulation of the interference membrane 709 may be customizedto allow for restricting or denying the passage of certain molecules tounderlying layers, for example, restricting or denying the passage oflarge molecules or of particular target molecules.

Interference membrane 709 is a solid coating surrounding the platinumwire 710. In this way, the expense and uncertainty of providing a windowthrough an insulating layer is avoided. Accordingly, the interferencemembrane 709 may be precisely coated or deposited over the platinum wire710 in a way that has a predictable and consistent passage of hydrogenperoxide. Further, the allowable area of interaction between thehydrogen peroxide and the surface of the platinum wire 710 isdramatically increased, as the interaction may occur anyplace along theplatinum wire 710. In this way, the interference membrane 709 enablesand increased level of interaction between the hydrogen peroxidemolecules in the surface of the platinum wire 710 such that theproduction of electrons is substantially amplified over prior artworking electrodes. In this way, the interference membrane enables thesensor to operate at a higher electron current, reducing the senor'ssusceptibility to noise and interference from contaminants, and furtherenabling the use of less sophisticated and less precise electronics inthe housing. In one non-limiting example, the ability to operate at ahigher electron flow allows the sensor's electronics to use morestandard operational amplifiers (op-amp), rather than the expensiveprecision op-amps required for prior art sensor systems. The resultingimproved signal to noise ratio allows enable simplified filtering aswell as streamlined calibration.

Further, during the manufacturing process it is possible to removeoxidation on the outer surfaces of the platinum wire 710 prior todepositing the interference membrane 709. Since the interferencemembrane 709 acts to seal the platinum wire 710, the level of oxidationcan be dramatically reduced, again allowing for a larger interactionsurface and further amplification of the glucose signal, resulting inhigher electron flow and enabling a higher signal to noise ratio. Inthis way, the new interference layer prevents fouling of the platinum'selectrical interface by eliminating undesirable oxidative effects.

In some embodiments, the interference membrane 709 is nonconductive ofelectrons, but is conductive of ions. In practice, a particularlyeffective interference membrane may be constructed using, for example,Poly-Ortho-Aminophenol (PoAP). PoAP may be deposited onto the platinumwire 710 using an electrodeposition process, at a thickness that can beprecisely controlled to enable a predictable level of hydrogen peroxideto pass through the interference membrane 709 to the platinum electrode710. Further, the pH level of the PoAP may be adjusted to set adesirable permselectivity for the interference membrane 709. Forexample, the pH may be advantageously adjusted to significantly blockthe passage of larger molecules such as acetaminophen, thereby reducingcontaminants that can reach the platinum wire 710. It will be understoodthat other materials may be used, for example, polyaniline, naphthol orpolyethelenediamine.

Sensor 700 also has a reference electrode 705 separate from workingelectrode 703. In this way, the manufacture of the working electrode issimplified and can be performed with a consistency that contributes todramatically improved stability and performance. The reference electrode705 is constructed of silver or silver chloride 714.

Referring now to FIG. 7B, another sensor 701 for a continuous biologicalmonitor is illustrated. Sensor 701 is similar to sensor 700, so will notbe described in detail. Sensor 701 has the same working electrode 703 asdescribed with reference to sensor 700. However, sensor 701 has areference wire 725 that has a silver/silver chloride layer 726surrounded by an ion limiting membrane 728. The application of this ionlimiting membrane 728 over the silver/silver chloride layer 726desirably controls the current sensitivity of the overall sensor device701 by controlling the flow of ions from the silver/silver chloridelayer 726. In this way, current sensitivity may be advantageouslycontrolled and defined. As will be understood, this can also act as asecondary method to control sensor sensitivity by controlling thechloride release from the electrode's surface.

Referring now to FIG. 8, a general description of a process 800 forformulating and applying the interference membrane is illustrated. Asshown in step 802, a conductive substrate is provided. This conductivesubstrate may be in the form of an elongated wire, but it will beappreciated that the conductive substrate can be provided in otherforms, such as printed or in the form of conductive pads. In someembodiments, the conductive substrate is a solid platinum wire, a lessexpensive wire that has been coated with platinum, or as disclosedherein, the conductive substrate may be a conductive carbon compoundcoated on a plastic substrate. It will be appreciated that otherconductive substrates may be used.

As shown in step 804, the interference membrane compound is nowprepared. This compound is formulated to be 1) non-electricallyconducting; 2) ion passing; and 3) permselective. Further, the compoundis particularly formulated to be electrodeposited in a thin and uniformlayer, and that has a thickness that is self-limiting due to the natureof electrically driven cross-linking. In this way, the compound may beapplied in a way that provides a well-controlled regulation of hydrogenperoxide molecule passage using a simple and cost-effectivemanufacturing processes. Further, the passage of the hydrogen peroxidecan occur over a much larger surface area as compared to prior artworking wires.

Generally, the characteristics of the present interference membranesidentified above can be formulated by mixing a monomer with a mildlybasic buffer, and converting the monomer into a more stable and usablepolymer by applying an electropolymerization process. In oneformulation:

a) Monomer: e.g., 2-Aminophenol, 3-Aminophenol, 4-Aminophenol, Aniline,Naphthol, phenylenediamine, or blends thereof.

b) Buffer: e.g., Phosphate Buffered Saline (PBS) tuned to about 7.5 toabout 10 pH, such as 7.5 to 9 pH, such as 8 pH by adding SodiumHydroxide.

c) Mix the monomer and buffer and electropolymerize.

d) Create a polymer; e.g., Poly-Ortho-Aminophenol (PoAP).

In the particular formulation set out above, the 2-Aminophenol monomeris mixed with a PBS buffer being mildly basic at a pH 8. The pH of thePBS buffer is adjusted using an additive, such as Sodium Hydroxide. Itwill be understood that the pH may be adjusted to create alternativeformulations consistent with this disclosure. For example, the pH of thecompound may be adjusted such that the permselectivity of the resultingPoAP can be modified. More particularly, PoAp may be formulated to havea defined molecular weight cutoff. That is, by adjusting the pH of theformulation, the PoAP may be modified to substantially restrict thepassage of molecules having a molecular weight larger than the cutoffmolecular weight. Accordingly, the PoAP can be modified according to themolecular weight of the contaminants that need to be restricted fromreaching the platinum wire. It will also be understood that othermonomers may be selected, and these alternative monomers may provide thedesired functional characteristics at a different pH. The 2-Aminophenoland PBS mixture is electropolymerized into Poly-Ortho-Aminophenol(PoAP.)

Optionally, the oxides or oxide layers may be removed from the surfaceof the conductive platinum substrate as illustrated in block 805. Asdescribed earlier, these oxides or layer of oxides dramatically restrictthe surface area available to the hydrogen peroxide to react with theplatinum. By removing these oxides or oxide layers, for example bychemical etching or physical buffing, a less contaminated platinum wiremay be provided for coating. In this way, the surface area of platinumavailable for hydrogen peroxide interaction is dramatically increased,thereby increasing the overall electrical sensitivity of the sensor.

The interference compound is then applied to the conductive substrate asshown in block 807. In one particular application, the interferencecompound is electrodeposited onto the conductive substrate, whichdeposits the compound in a thin and uniform layer. Further, theelectrodeposition process facilitates a chemical cross-linking of thepolymers as the PoAP is deposited. It will be understood that otherprocesses may be used to apply the polymer to the conductive substrate.

As described above, the interference membrane has a compound that isself-limiting in thickness. The overall allowable thickness for themembrane may be adjusted according to the ratio between the monomer andthe buffer, as well as the particular electrical characteristics usedfor the electropolymerization process. Also, the interference membranemay be formulated for a particular permselective characteristic byadjusting the pH. It will also be understood that a cyclic voltammetry(CV) process may be used to electrodeposit the interference membranecompound, such as PoAP. A CV process is generally defined by having (1)a scanning window that has a lower voltage limit and upper voltagelimit, (2) a starting point and direction within that scanning window,(3) an elapsed time for each cycle, and (4) the number of cyclescompleted. It will be understood by one skilled in the art that thesefour factors can provide nearly infinite alternatives in the preciseapplication of the interference membrane compound. In one example, thefollowing ranges have been found to be effective for the CV process toapply PoAP:

-   Scanning window: −1.0V to 2.0V-   Starting point: −0.5V to 0.5 VV-   Rate: x-y cycles per minute-   Cycles 5-50

As illustrated in step 811, the enzyme layer is then applied, whichincludes the glucose oxidase, and then a glucose limiting layer isapplied as shown in 818. This glucose limiting layer, as discussedabove, is useful to limit the number of glucose molecules that areallowed to pass into the enzyme layer.

Finally, as illustrated in block 821, an insulator may be applied to thereference wire. In many cases, the reference wire will be asilver/silver oxide wire, and the insulator will be an ion limitinglayer that is nonconductive of electrons.

Glucose-limiting Layer

Referring now to FIG. 9A, a sensor 900 is illustrated for use with acontinuous biological monitor. Sensor 900 has a working electrode 903and a reference electrode 905. The silver/silver chloride referenceelectrode 914, the conductive layer 910, the interference membrane 909,and enzyme layer 908 are similar to those as discussed previously withreference to sensor 700, so will not be discussed in detail. It will beunderstood that several alternatives exist for these layers consistentwith this disclosure.

Sensor 900 has a glucose limiting membrane layer 907. As will bedescribed, the glucose limiting membrane 907 may be manufactured usingsimple and inexpensive manufacturing techniques, and provides a moreuniform glucose limiting membrane with more precise regulation for theglucose molecules. In this way, the level of glucose molecules permittedto pass into the enzyme layer may be more precisely and uniformlydefined and controlled, and resulting calculations and results areenabled to be more linear and precise. The glucose limiting membrane 907is constructed to provide a thin conformal layer of a physicallycross-linked material that is easy to dispose and that providesexceptional uniformity, glucose molecule control, and linearity results.In one specific example, the physically cross-linked material useshydrogen-bonds. Importantly, the glucose limiting layer does not relyupon chemical cross-linking.

Although the present glucose limiting layer 907 is illustrated withsensor 900, it will be appreciated that the glucose limiting layer 907may also be advantageously used on other sensors, such as prior artsensor 600. It will be appreciated that the present glucose limitinglayer may be broadly used on other types of biological sensors.

As formulated in FIG. 9A, the glucose limiting layer 907 may beformulated to provide a uniform layer that more evenly and preciselypasses glucose molecules into the enzyme layer 908 than typical priorart glucose limiting layers, which results in a more stable, consistent,and accurate generation of free electrons. As the sensor 900 more evenlyand uniformly passes glucose and generates electrons, the sensor is moreaccurate, has less sensitivity to noise, is more stable, and is morereadily calibrated.

Referring now to FIG. 9B, another sensor 901 for a continuous biologicalmonitor is illustrated. Sensor 901 is similar to sensor 900, so will notbe described in detail. Sensor 901 has the same working electrode 903 asdescribed with reference to sensor 900. However, sensor 901 has areference wire 925 that has a silver/silver chloride layer 926surrounded by an ion limiting membrane 928. The application of this ionlimiting membrane 928 over the silver/silver chloride layer 926desirably controls the current sensitivity of the overall sensor device901 by controlling the flow of ions from the silver/silver chloridelayer. In this way, current sensitivity may be advantageously controlledand defined.

Referring now to FIG. 10, a process 1000 for creating a glucose limitinglayer is generally described. Before providing details and examples, theprocess is generally described. First, a hydrophilic bonding material isselected as shown in step 1002. Also, a hydrophobic bonding material isselected as shown in step 1004, and a solvent is selected as illustratedin step 1007. The hydrophilic bonding material, the hydrophobic bondingmaterial, and the solvent are mixed together in a desired ratio, whichresults in a bonding gel as illustrated in step 1011. This bonding gelmay then be applied over the enzyme layer 1018 on the working wire. Thegel then cures to form strong and resilient hydrogen-bonded structuresas illustrated at 1021. It will be understood that other materials maybe used, and that other types of physical crosslinking may be formed.

In selecting the hydrophilic bonding material 1002, it is desirable toidentify a hydrophilic bonding material that has a relatively highmolecular weight, for example 1 to 5 million. It has been found thathydrophilic bonding material with a molecular weight of 1 to 3 millionis particularly effective. As understood, the molecular weight of apolymer is the sum of the atomic weights of all the atoms in themolecule. Accordingly, the selected hydrophilic bonding material istypically a significantly large polymer. Further, the hydrophilic bodymaterial is selected to be readily dispensable in standard manufacturingprocesses, and to have the ability to form strong hydrogen bonds.Although in some embodiments the hydrophilic bonding material has arelatively high molecular weight, is readily dispensable, and has stronghydrogen bonds capability, it will be understood that othercharacteristics may become important depending upon the specificapplication. For example, Polyvinyl alcohol, Polyacrylic acid, orPolyvinylpyrrolidone (PVP) may be used as a hydrophilic bonding materialfor the glucose limiting layer. In one specific example, PVP, in itspharma grade form, has a molecular weight of about 1.3 million. It willbe understood that other polymers may be found with similar or otherdesirable characteristics.

The hydrophobic bond material is then selected 1004. In particular, thehydrophobic material is selected based on a desirable biocompatibility,as well as a ratio between hard and soft segments. Generally,hydrophobic materials are formed of segments of small monomers which arecross-linked to much larger polymer sections. A higher proportion ofsoft segments allows the hydrophobic bond material to have a higherdegree of interaction with the solvent and hydrophilic bonding material;however, the higher ratio of soft segments also decreases thehydrophobic characteristic of the material, as the small segments tendto be hydrophilic. As to the hard segments, a higher ratio of hardsegments provides for a stronger physical characteristic, which is oftenmeasured as a Shore hardness using a durometer. In this way, ahydrophobic material may be selected that has an appropriate level ofinteraction with the solvent and hydrophilic materials, as well ashaving sufficient hardness to act effectively as a protective coating.In some embodiments, polyurethane may be used as a hydrophobic bondmaterial, with the desired characteristics of both providing sufficienthardness, as well as desirable interaction with the hydrophilic bondmaterial (e.g., PVP) and the selected solvent. Additionally, siliconesmay also be used as the hydrophobic bonding material. It will also beappreciated that other types of hydrophobic bonding materials may beselected according to application-specific needs

The third material in step 1007 is the solvent. Generally, a solvent isselected that is polar, binary, and sufficiently volatile for the curingneeds. First, the solvent should have sufficiently strong polarcharacteristics to assist in properly aligning the hydrophilic andhydrophobic bonding materials. Second, as the solvent must dissolve boththe hydrophilic and the hydrophobic material, the solvent should beselected for advantageous dissolving characteristics for each of theselected bonding materials. It will be understood that there are trinarysolvents that can be substituted. Finally, the volatility of the solventshould be selected to support the desired cure characteristics. Forexample, some applications may need to be completed in a short period oftime, thus requiring a fast flash solvent. In other cases, less volatilesolvents may be substituted. In one example, a mixture of a heavyorganic compound with an alcohol may provide a desirable solvent for theglucose limiting layer. In a specific example, the heavy organiccompound may be tetrahydrofuran (THF) or Dimethylformamide (DMF), andthe alcohol may be ethanol. It will be understood that other compoundsmay be used that can provide the desirable characteristics of thesolvent.

The hydrophilic bonding material, hydrophobic bond material, and solventare mixed together to form a bonding gel in step 1011. The viscosity ofthe bonding gel may be tuned by adjusting the ratio of the solvent tothe bonding materials. The bonding gel can then be applied over theenzyme layer in step 1018. The bonding gel is easy to work with, and maybe dipped, sprayed, deposited or pad printed using various manufacturingprocesses. The bonding gel is then cured in step 1021, which may be donein ambient air, by using added heat, or by using added vacuum. It willbe understood that other processes may be used to either speed or slowthe curing process. As the bonding material cures, the hydrophobic andhydrophilic physically cross-link, and in particular, form hydrogenbonds. The resulting hydrogen bonded layer enables a highly desirableuniform and even passage of glucose molecules as compared to priorchemically bonded layers.

Enzyme Layer

As discussed with reference to sensors 600, 700, and 900, the workingwire for each sensor has a respective enzyme layer 608, 708 and 908. Asis well known, the enzyme layer facilitates a chemical interactionbetween glucose and glucose oxidase (GOx), which generates hydrogenperoxide (H₂O₂). The hydrogen peroxide further reacts with a conductiveplatinum substrate, which generates a current of free electrons that canbe measured, with the measured level of current being proportional tothe level of glucose in the bloodstream or in another bodily fluid, suchas ISF. To make a useful membrane for a glucose sensor, GOx is oftenstabilized with glutaraldehyde, imidoesters (dimethyl adipimidate,dimethyl suberimidate), hydroxysuccinimide and its derivatives.Typically, a formulation of about 0.6% glutaraldehyde is mixed with theGOx, and the mixture is then applied to the working wire. It will beunderstood that other ratios may be used, and that there may be otheradditives in the mixture. It is well known that polyaziridine may alsobe used as a stabilizer for GOx, but it suffers from similardisadvantages as glutaraldehyde.

Unfortunately, even with proper mixing, the GOx does not disperse evenlywithin the glutaraldehyde, leaving portions of the enzyme layer with ahigher concentration of GOx and portions with a lower concentration ofGOx. This uneven distribution of GOx causes a non-uniform interactionbetween glucose and GOx, which leads to an uneven generation of hydrogenperoxide. That is, given a constant level of glucose, different portionsof the enzyme layer will generate more or fewer hydrogen peroxidemolecules, resulting in more or fewer free electrons. In this way, themeasured glucose level may vary based on where on the enzyme layer theglucose molecule lands and reacts. This uncertainty and variability dueto uneven GOx dispersion may lead to an erroneous blood sugar reading.

Further, commercial GOx is derived from bacterial or fungal sources andas such is known to be cytotoxic, that is, harmful to cells. Even whenGOx is stabilized with glutaraldehyde, some GOx may move within thelayer and leech from the enzyme layer into the subject's body. Sincenone of the known protective layers for an enzyme layer can fully entrapGOx, there is a risk that at least some GOx may be exposed to thesubject's cells.

To address the deficiencies in the known glutaraldehyde-stabilized GOx,embodiments of an enzyme layer are provided that provide forsubstantially improved GOx entrapment and even distribution. As shown inFIG. 11, the enzyme layer is made using a process 1100. An aqueousemulsion of polyurethane is made as shown at step 1103. It will beunderstood that the amount of water that is mixed with the polyurethanemay be adjusted according to application-specific requirements. Althoughpolyurethane shall be used in the description of FIG. 11, it will beunderstood that other emulsions may be substituted, such as aqueoussilicone dispersions. As illustrated at step 1105, the aqueouspolyurethane emulsion is mixed with an aqueous acrylic polyol emulsion.The acrylic polyol acts as a self cross-linker with the polyurethane togenerate a highly stable and tight structure that is able to fullyentrap the GOx. The combination of the polyurethane emulsion from step1103 and the acrylic polyol emulsion from step 1105 generates a baseemulsion in step 1107. Depending upon application-specific requirements,the ratio of polyurethane to acrylic polyol may be adjusted; however, inone example approximately equal amounts of each are mixed together toform the base emulsion in step 1107. In some embodiments, GOx is blendedwith the polyurethane at a ratio of about 1 part GOx to 60 partspolyurethane by volume. It will be understood that other ratios may beused depending upon the particular application. It will also beunderstood that if other metabolic functions are to be tested other thanglucose levels, other enzymes may be used.

As shown in block 1110 other optional additives may be added. Forexample, one or more hydrophiles may be added to the emulsion mixture tofacilitate better mixing or to provide a more appropriate viscosity forapplication. Examples of hydrophiles that may be used in the formulationof the enzyme layer include PVP, PEO and Si-PEO. Si-PEO is understood toinclude silanes and PDMS PEO. It will be understood that otherhydrophiles may be used. Although the acrylic polyol can provide forself cross-linking as it cures, other cross-linking polymers may beadded for additional cross-linking. For GOx, such cross-linkers mayinclude, for example, glutaraldehyde, imidoesters, hydroxysuccinimide,carbodilite, melamines, epoxies and polyaziridines.

The polyurethane/GOx blend is applied to the working electrode in step1109, for example, by spraying, dipping, depositing or printing, 1121.The blend is cured in step 1125, at which time the layer cross-links toprovide a stable dispersion of GOx.

Advantageously, the polyurethane/GOx blend, being an aqueous emulsion,is safe, easy to handle and apply, and provides for an even distributionof GOx. Further, as the cross-linked polymers are fully stabilized andentrapped within the layer, GOx is not able to move from the enzymelayer to the subject's body, eliminating safety concerns. Also, as thepolyurethane/GOx blend is more stable than prior art enzyme layers, ithas a longer usable shelf life, and exhibits the ability to support ahigher loading. With a higher loading of GOx, the polyurethane/GOx blendhas increased sensitivity and enables higher signal to noise ratios.

Sensors

Referring now to FIG. 12, a sensor 1200 according to some embodiments isillustrated. Sensor 1200 is constructed with a working electrode 1203and a reference electrode 1205. As described previously, referenceelectrode 1205 is typically silver chloride or silver 1214. Workingelectrode 1203 has a glucose limiting layer 1207 as described withreference to FIG. 9A, FIG. 9B and FIG. 10. Working electrode 1203 alsohas an enzyme layer 1208 in contact with the glucose limiting layer 1207as described with reference to FIG. 11. The enzyme layer 1208 is also incontact with an interference membrane 1209 as described with referenceto FIG. 7A, FIG. 7B and FIG. 8. The interference membrane 1209 issupported by substrate 1210, as fully described with reference to FIGS.4A-C and FIG. 5. As illustrated, substrate 1210 has a plastic substrateportion 1210 b that has a carbon coating 1210 a, as described withreference to FIGS. 4A-C and FIG. 5. It will be understood, asillustrated, that the carbon coating 1210 a will include a catalyst forhydrogen peroxide such as Phthalocyanine or Prussian blue.

Advantageously, as compared to prior sensors, sensor 700 provides anelectrical signal with a higher signal-to-noise ratio, is less expensiveto manufacture, and is safer for a patient to wear.

Enzyme Layer Having Direct Electron or Peroxide Generation

Referring now to FIG. 13A, a sensor 1300 is illustrated. Sensor 1300 isillustrated as a two wire sensor having a working electrode 1303 and areference electrode 1305. As described with previous sensors, thereference electrode 1305 is generally silver chloride or silver 1314. Itwill be understood that the construction of sensor 1300 may be any ofthe common structures for two wire sensors.

The working wire 1303 has a glucose limiting layer 1307. The glucoselimiting layer may be of a known construction, but as illustrated, theglucose limiting layer 1307 is the glucose limiting layer described withreference to FIG. 9A and FIG. 9B. As described earlier, a glucoselimiting layer is provided to limit and control the number of glucosemolecules that may be passed from a patient's blood or ISF into theenzyme layer, thereby improving linearity of the overall sensorresponse. The glucose limiting layer still allows oxygen to travel tothe enzyme layer. An interference membrane 1309 may be positioned belowthe glucose limiting layer 1307. In one example, the interferencemembrane 1309 is the interference membrane described with reference toFIG. 7A and FIG. 7B. As such, the interference layer 1309 ispermselective to reject the passage of larger molecules. In this way,large molecules such as acetaminophen or other contaminants may beblocked from reaching the enzyme layer.

Sensor 1300 has a plastic material substrate 1312, such as a plasticwire. This plastic substrate material is selected to have sufficientstrength for being inserted under the skin of a patient, as well asflexibility for patient comfort and ease of manufacturing. Further, itwill be understood that the plastic substrate should be biologicallysafe and generally electrically unreactive. It will be understood that awide range of materials meet the mechanical and functional requirementsfor the selected plastic substrate. For example, numerous organicpolymers and thermoplastics may be used. For illustrative purposes only,the following specific plastic substrate materials may be used:polyethylene, polypropylene, polystyrene, polyvinyl chloride, andpolylactic acid. It will be appreciated that a wide variety of materialsmay be used as the plastic substrate. This plastic substrate can beformed into an elongated wire in many shapes to support construction ofdifferent types of sensors. Generally, these plastic wires would beformed using well-known extrusion processes.

The plastic substrate 1312 supports a carbon-enzyme layer 1310.Generally, the new carbon-enzyme layer 1310 is prepared as a coatingthat is an aqueous dispersion of carbon materials, an elastomericmaterial, cross-linkers and GOx. For example, the carbon material may bein the form of graphite, graphene, diamagnetic graphite, pyrolyticcarbon, carbon black, carbon paste, or carbon ink. In some cases, tosupport particular applications, other additives may be added to thecarbon compound for enhanced electrical and response characteristics.

There are several acceptable elastomeric materials that would providethe desired characteristics, for example: polyurethane, silicone,acrylates or acrylics. It will be understood that other elastomericmaterials may be substituted. Also, in some embodiments the carboncompound, after the elastomeric material is cured, does not delaminatefrom the plastic wire, in contrast to platinum coating of a Tantalumwire. In one example, an aqueous polyurethane emulsion is selected to bemixed with an aqueous acrylic polyol dispersion as a cross-linker. Itwill be understood that alternative or additional cross-linkers andother additives may be used.

As described with reference to sensor 1300, the GOx enzyme is used, assensor 1300 is directed to detecting glucose levels. It will beunderstood that other enzymes such as lactate dehydrogenase (lactate),hydroxybutyrate dehydrogenase (ketone) may be used for other metabolicsensors. It will also be understood that if an enzyme other than GOx isused, additional modifications may be needed in the selection and ratiosof materials in the carbon-enzyme layer.

The carbon-enzyme coating 1310 is applied to the plastic substrate 1312and cured. As the aqueous dispersion cures on the plastic substrate, itcrosslinks to a flexible but strong coating for the substrate, with theGOx evenly dispersed and fully entrapped. Further, the selectedcombination of the carbon materials provides for advantageousstructural, mechanical, and electrical properties of the carbon-enzymelayer.

In operation, sensor 1300 permits glucose and oxygen to pass through theglucose limiting layer 1307 and the interference membrane 1309 into thecarbon-enzyme layer 1310. The interference membrane 1309 blocks largermolecules that may contaminate or interfere with the chemical andelectrical processes. Once the glucose and oxygen pass into thecarbon-enzyme layer 1310, they react to form hydrogen peroxide, whichthen interacts with the carbon to generate free electrons. These freeelectrons may then be conducted as shown by arrow 1315 through thecarbon-enzyme layer 1310 to the electronics of the sensor 1300.

Advantageously, sensor 1300 does not use any platinum, and has the GOxenzyme evenly dispersed and fully trapped within the carbon-enzymelayer. In this way, the sensor 1300 reduces the risk of potential safetyissues with GOx, and provides a highly desirable level of sensitivityand high signal-to-noise performance.

Referring to FIG. 13B, sensor 1301 is illustrated. Sensor 1301 issimilar to sensor 1300, so will not be described in detail. Sensor 1301is a single wire sensor, where reference electrode 1305 is attached toworking electrode 1303. It will be understood that such a physicalconstruction may be achieved through various printing, extrusion anddeposition processes.

Referring now to FIG. 14A, sensor 1400 is illustrated. Sensor 1400 issimilar to sensor 1300, so will not be described in detail. Asillustrated, sensor 1400 has a glucose limiting layer 1407 similar toglucose limiting layer 1307, a carbon-enzyme layer 1410 similar tocarbon-enzyme layer 1310, plastic substrate 1412 similar to plasticsubstrate 1312, and reference electrode 1405 similar to referenceelectrode 1305, so none of these will be described in detail. Asillustrated, sensor 1400 does not have an interference membrane. In somecases, an interference membrane will not be needed, as one of theprimary purposes for the interference membrane is to protect largemolecule contaminants from reaching and contaminating the platinum wire.Since sensor 1300 uses no platinum wire, the need for an interferencemembrane may be reduced.

Referring to FIG. 14B, sensor 1401 is illustrated. Sensor 1401 issimilar to sensor 1400, so will not be described in detail. Sensor 1401is a single wire sensor, where reference electrode 1405 with silverchloride or silver 1414 is attached to working electrode 1403. It willbe understood that such a physical construction may be achieved throughvarious printing and deposition processes.

Referring now to FIG. 15, single wire sensor 1501 is illustrated. Sensor1501 has a working wire 1505 that is physically attached to a referencewire 1507, where both working wire 1505 and reference wire 1507 havesemi-circular cross-sections with flat surfaces facing each other. Insome cases, an insulating member 1509 may be positioned at the flatsurface interface between the working wire 1505 and the reference wire1507. In some embodiments, the working wire 1505 may be the working wiredescribed with reference FIG. 13B or FIG. 14B. Still referring to FIG.15, another single wire sensor 1511 is illustrated. Single wire sensor1511 has a reference wire 1515 attached to an insulated substrate 1513that has a triangular cross-section. In one example, the substrate 1513may be an extruded plastic wire. Also attached to the substrate 1513 isa reference electrode 1517 and a counter electrode 1519. In someembodiments, the working wire 1515 may be the working wire describedwith reference FIG. 13B or FIG. 14B. The reference wire 1515, referenceelectrode 1517 and counter electrode 1519 have flat surfaces that facethe substrate 1513.

Referring now to FIG. 16, a process 1600 for making and applying thecarbon-enzyme layer as shown in FIG. 13A, FIG. 13B, FIG. 14A and FIG.14B is described. As discussed with reference to sensors 600, 700, and900, the working wire for each sensor has a respective enzyme layer 608,708 and 908. As is well known, the enzyme layer facilitates a chemicalinteraction between glucose and glucose oxidase (GOx), which generateshydrogen peroxide (H₂O₂). The hydrogen peroxide further reacts with aconductive platinum substrate, which generates a current of freeelectrons that can be measured, with the measured level of current beingproportional to the level of glucose in the bloodstream or ISF. To makea useful membrane for a glucose sensor, GOx is often stabilized withglutaraldehyde. Typically, a formulation of about 0.6% glutaraldehyde ismixed with the GOx, and the mixture is then applied to the working wire.It will be understood that other ratios may be used, and that there maybe other additives in the mixture. It is well known that polyaziridine,imidoesters, hydroxysuccinimide, carbodilite, melamines, epoxies,benzoyl peroxide, dicumyl peroxide may also be used as a stabilizer forGOx, but it suffers from similar disadvantages as glutaraldehyde.

Unfortunately, even with proper mixing, the GOx does not disperse evenlywithin the glutaraldehyde, leaving portions of the enzyme layer with ahigher concentration of GOx and portions with a lower concentration ofGOx. This uneven distribution of GOx causes a non-uniform interactionbetween glucose and GOx, which leads to an uneven generation of hydrogenperoxide. That is, given a constant level of glucose, different portionsof the enzyme layer will generate more or fewer hydrogen peroxidemolecules, resulting in more or fewer free electrons. In this way, themeasured glucose level may vary based on where on the enzyme layer theglucose molecule lands and reacts. This uncertainty and variability dueto uneven GOx dispersion may lead to an erroneous blood sugar reading.

Further, GOx is known to be cytotoxic, that is, harmful to cells. Evenwhen GOx is stabilized with glutaraldehyde, some GOx may move within thelayer and leech from the enzyme layer into the subject's body. Sincenone of the known protective layers for an enzyme layer can fully entrapGOx, there is a risk that at least some GOx may be exposed to thesubject's cells.

Additionally, most known sensors use substrate of either solid platinumor a platinum coated metal wire. Either way, platinum is expensive, andleads to a higher priced sensor. The platinum is also susceptible tooxidation, which can lead to instability and a low signal-to-noiseratio. As discussed with reference to FIG. 4A, FIG. 4B, FIG. 4C and FIG.5, a carbon compound may be applied over a plastic substrate, removingthe need for any platinum in the working wire. As was described withreference to process 500, the carbon coating can be constructed to beflexible, strong, with desirable electrical characteristics, and appliedin a manner that it will not delaminate from the plastic substrate.

To address the deficiencies in the known glutaraldehyde-stabilized GOx,and the expense of platinum, a new carbon-enzyme layer is provided thatprovides for substantially improved GOx entrapment and evendistribution, with carbon in the same layer to provide strength,flexibility, and appropriate electrical characteristics. Further, thecarbon-enzyme layer has been found to provide for a direct generation offree electrons, not only eliminating the need for any ion conductinglayer or platinum wire, but providing for enhanced stability andsubstantially improved signal-to-noise ratios.

As shown in FIG. 16, the carbon-enzyme layer is made using a process1600. An aqueous emulsion of polyurethane is made as shown at step 1603.It will be understood that the amount of water that is mixed with thepolyurethane may be adjusted according to application-specificrequirements. Although polyurethane has proven to perform well, it willbe understood that other emulsions may be substituted, such as aqueoussilicone dispersions As illustrated at step 1605, the aqueouspolyurethane emulsion is mixed with an aqueous acrylic polyol emulsion.The acrylic polyol acts as a self cross-linker with the polyurethane togenerate a highly stable and tight structure that is able to fullyentrap the GOx. The combination of the polyurethane emulsion in step1603 and the acrylic polyol emulsion in step 1605 generates a baseemulsion in step 1607. Depending upon application-specific requirements,the ratio of polyurethane to acrylic polyol may be adjusted; however, inone example approximately equal amounts of each are mixed together toform the base emulsion in step 1607. In one embodiment, GOx is blendedwith the polyurethane at a ratio of about 1 part GOx to 60 partspolyurethane by volume. It will be understood that other ratios may beused depending upon the particular application. It will also beunderstood that if other metabolic functions are to be tested other thanglucose levels, other enzymes may be used such as ketone, lactate, orother metabolic catalysts in step 1609.

As shown in block 1610 other optional additives may be added. Forexample, one or more hydrophiles may be added to emulsion mixture tofacilitate better mixing or to provide a more appropriate viscosity forapplication. Well known hydrophiles include PVP, PEO and Si-PEO. Si-PEOis understood to include silanes and PDMS PEO. It will be understoodthat other hydrophiles may be used. Although the acrylic polyol canprovide for self cross-linking as it cures, other cross-linking polymersmay be added for additional cross-linking. For GOx, such cross-linkersmay include, for example glutaraldehyde and polyaziridine.

A blend of carbon material is also mixed into the emulsion in step 1612.The carbon materials and ratios are selected according to applicationand functional requirements. For example, carbon (graphite) may be addedto increase strength of the resulting layer, while graphene, pyrolyticgraphite or a blend of graphene and pyrolytic graphite may be added toprovide for improved electrical characteristics. The ratio of graphiteto be added will be selected to provide sufficient strength for theresulting carbon-enzyme layer, but still permit sufficient flexibilityin the layer so as not to delaminate from the plastic substrate or be sobrittle as to crack. Also, the amount of graphene and pyrolytic graphitecan be adjusted to set a desirable resistance for the carbon-enzymelayer. It will be understood that other forms of carbon may besubstituted.

As the final mixture is prepared in step 1612, additional water or oneor more hydrophiles (such as PVP) may be added in step 1616 to obtainthe proper viscosity and fluid properties (e.g., to thin the mixture) tofacilitate both even dispersion of the GOx, and to enable the selectedapplication technique.

While a particular order for adding components of the polyurethane/GOx/Cblend in step 1612 has been illustrated, it will be understood that theorder may be changed without affecting the resulting layer. Thepolyurethane/GOx/C blend is applied to the working electrode, forexample, by spraying, dipping, depositing or printing in step 1621. Theblend is cured in step 1625, at which time the layer cross-links toprovide a stable dispersion of GOx.

Advantageously, the polyurethane/GOx/C blend, being an aqueous emulsion,is safe, easy to handle and apply, and provides for an even distributionof GOx. Further, as the cross-linked polymers are fully stabilized andentrapped within the layer, GOx is not able to move from the enzymelayer to the subject's body, reducing safety concerns. Also, as thepolyurethane/GOx/C blend is more stable than prior enzyme layers, it hasa longer usable shelf life, and exhibits the ability to support a higherloading. With a higher loading of GOx, the polyurethane/GOx/C blend hasincreased sensitivity and enables higher signal to noise ratios thanknown devices.

In one example of the applied carbon-enzyme emulsion in step 1621, thecarbon-enzyme emulsion comprises the following:

0.5 to 2 parts polyurethane emulsion;

0.5 to 2 parts acrylic polyol emulsion;

0.5 to 2 parts carbon, comprising:

-   -   0.5 to 1 parts graphite;    -   0.0 to 1 parts graphene; and    -   0.0 to 2 parts pyrolytic graphite;

0.0 to 3 parts water and hydrophile; and

0.01 to 0.1 parts GOx.

Although process 1600 has been discussed with reference to the GOxenzyme, it will be appreciated that other enzymes may be substitutedaccording to the particular metabolic function to be monitored. Forexample, the following enzymes may be used in process 1600. It will beappreciated that other enzymes and metabolic functions can be used, forexample:

Enzyme Metabolic Function Lactate dehydrogenase Lactate Hydroxybutyratedehydrogenase Ketone

Reference has been made in detail to embodiments of the disclosedinvention, one or more examples of which have been illustrated in theaccompanying figures. Each example has been provided by way ofexplanation of the present technology, not as a limitation of thepresent technology. In fact, while the specification has been describedin detail with respect to specific embodiments of the invention, it willbe appreciated that those skilled in the art, upon attaining anunderstanding of the foregoing, may readily conceive of alterations to,variations of, and equivalents to these embodiments. For instance,features illustrated or described as part of one embodiment may be usedwith another embodiment to yield a still further embodiment. Thus, it isintended that the present subject matter covers all such modificationsand variations within the scope of the appended claims and theirequivalents. These and other modifications and variations to the presentinvention may be practiced by those of ordinary skill in the art,without departing from the scope of the present invention, which is moreparticularly set forth in the appended claims. Furthermore, those ofordinary skill in the art will appreciate that the foregoing descriptionis by way of example only, and is not intended to limit the invention.

What is claimed is:
 1. A working electrode for a subcutaneous sensor foruse with a continuous biological monitor for a patient, comprising: aconductive substrate; and a carbon-enzyme layer on the conductivesubstrate, the carbon-enzyme layer comprising: a polyurethanecrosslinked with an acrylic polyol; an enzyme fully entrapped by thepolyurethane crosslinked with the acrylic polyol, the enzyme selectedaccording to a biological function to be monitored; and a carbonmaterial; wherein the carbon-enzyme layer is electrically conductive andfacilitates a generation of either peroxide or electrons within thecarbon-enzyme layer responsive to reacting the enzyme with a targetbiologic from blood of the patient.
 2. The working electrode accordingto claim 1, further including a glucose limiting layer over thecarbon-enzyme layer that limits an amount of glucose from the blood ofthe patient that passes to the carbon-enzyme layer.
 3. The workingelectrode according to claim 1, wherein the enzyme is glucose oxidase(GOx), lactate dehydrogenase, lactate oxidase, or hydroxybutyratedehydrogenase.
 4. The working electrode according to claim 1, whereinthe enzyme is glucose oxidase (GOx) and the biological function isglucose level.
 5. The working electrode according to claim 1, whereinthe polyurethane crosslinked with the acrylic polyol further comprises:an additional cross-linker selected from imidoesters,hydroxysuccinimides, carbodilites, melamines, epoxies, benzoyl peroxidesor dicumyl peroxides.
 6. The working electrode according to claim 1,wherein the carbon material comprises carbon, graphite, graphene, orpyrolytic graphite.
 7. The working electrode according to claim 1,further comprising a hydrophile selected from polyvinylpyrrolidone(PVP), polyethylene oxide (PEO) or silane-PEO (Si-PEO).
 8. The workingelectrode according to claim 1, wherein the carbon-enzyme layer has anapproximate ratio of materials comprising: 0.5 to 2.0 parts of thepolyurethane; 0.5 to 2.0 parts of the acrylic polyol; 0.01 to 0.1 partsof the enzyme; and 0.5 to 2.0 parts of the carbon material.
 9. Theworking electrode according to claim 8, wherein the carbon materialfurther comprises: 0.5 to 1.0 parts of a graphite; 0.0 to 1.0 parts of agraphene; and 0.0 to 2.0 parts of a pyrolytic graphite.
 10. The workingelectrode according to claim 8, wherein the carbon-enzyme layer furthercomprises 0.0 to 3.0 parts of water and a hydrophile.
 11. A workingelectrode for a subcutaneous sensor for use with a continuous biologicalmonitor for a patient, comprising: a conductive substrate; and acarbon-enzyme layer on the conductive substrate, the carbon-enzyme layercomprising: a silicone crosslinked with an acrylic polyol; an enzymefully entrapped by the silicone crosslinked with the acrylic polyol, theenzyme selected according to a biological function to be monitored; anda carbon material; wherein the carbon-enzyme layer is electricallyconductive and facilitates a generation of either peroxide or electronswithin the carbon-enzyme layer responsive to reacting the enzyme with atarget biologic from blood of the patient.
 12. The working electrodeaccording to claim 11, further including a glucose limiting layer overthe carbon-enzyme layer that limits an amount of glucose from the bloodof the patient that passes to the carbon-enzyme layer.
 13. The workingelectrode according to claim 11, wherein the enzyme is glucose oxidase(GOx), lactate dehydrogenase, lactate oxidase, or hydroxybutyratedehydrogenase.
 14. The working electrode according to claim 11, whereinthe enzyme is glucose oxidase (GOx) and the biological function isglucose level.
 15. The working electrode according to claim 11, whereinthe silicone crosslinked with the acrylic polyol further comprises: anadditional cross-linker selected from imidoesters, hydroxysuccinimides,carbodilites, melamines, epoxies, benzoyl peroxides or dicumylperoxides.
 16. The working electrode according to claim 11, wherein thecarbon material comprises carbon, graphite, graphene, or pyrolyticgraphite.
 17. The working electrode according to claim 11, furthercomprising a hydrophile selected from polyvinylpyrrolidone (PVP),polyethylene oxide (PEO) or silane-PEO (Si-PEO).
 18. The workingelectrode according to claim 11, wherein the carbon-enzyme layer has anapproximate ratio of materials comprising: 0.5 to 2.0 parts of thesilicone; 0.5 to 2.0 parts of the acrylic polyol; 0.01 to 0.1 parts ofthe enzyme; and 0.5 to 2.0 parts of the carbon material.
 19. The workingelectrode according to claim 18, wherein the carbon material furthercomprises: 0.5 to 1.0 parts of a graphite; 0.0 to 1.0 parts of agraphene; and 0.0 to 2.0 parts of a pyrolytic graphite.
 20. The workingelectrode according to claim 18, wherein the carbon-enzyme layer furthercomprises 0.0 to 3.0 parts of water and a hydrophile.